Magnetic apparatus for MRI

ABSTRACT

Magnetic apparatus for MRI/MRT probes and methods for construction thereof are disclosed. One embodiment includes a pair of opposed magnet assemblies defining an open region therebetween, a transmitting RF coil having at least a portion thereof disposed within the open region, at least one receiving RF coil disposed within the open region and X,Y and Z gradient coils. At least one of the X,Y and Z gradient coils is disposed outside of the open region. Another embodiment of the apparatus includes a single magnet assembly having a first surface and a second surface opposing the first surface, a transmitting RF coil having at least a portion thereof opposing the first surface, at least one receiving RF coil and X,Y and Z gradient coils. At least one of the X,Y and Z gradient coils opposes the second surface. In another embodiment the magnet assembly generates a permanent z-gradient magnetic field and therefore includes only X and Y gradient coils, at least one of which opposes the second surface. The apparatuses may also include one or more shim coils.

REFERENCE TO RELATED APPLICATIONS:

This application claims priority of and the benefit of U.S. provisionalapplication Ser. No. 60/059,659, filed Sep. 25, 1997.

FIELD OF THE INVENTION

The present invention is generally related to the fields of magneticresonance imaging (MRI) and magnetic resonance therapy (MRT).

BACKGROUND OR THE INVENTION

MRI systems for performing whole body imaging usually employ largemagnets which effectively surround the patient. Such magnets are usuallylarge superconductor magnets which are expensive and difficult tomaintain. MRI systems for performing local imaging of specific bodyparts or organs are known in the art.

U.S. patent application Ser. No. 08/898,773, now U.S. Pat. No. 5,900,793to Katznelson et al., filed Jul. 23, 1997 and entitled "PERMANENT MAGNETASSEMBLIES FOR USE IN MEDICAL APPLICATIONS", now U.S. Pat. No. 5,900,793and incorporated herein by reference discloses, inter alia, compactpermanent magnet assemblies for use in medical applications includingMRI and/or MRT.

A typical application using an intra-operative MRI system is brainsurgery. Reference is now made to FIG. 1 which is a schematicperspective view of a small organ dedicated MRI probe useful in brainsurgery. The MRI probe 1 includes two annular permanent magnetassemblies 2 and 4 connected by a frame 3. The frame 3 and the magnetassemblies 2 and 4 are shaped for imaging the brain of a patient 6.During MRI assisted brain surgery or MRT, the head of the patient 6 ispositioned between the two magnet assemblies 2 and 4.

Reference is now made to FIG. 2 which is a schematic isometric view ofthe two permanent magnet assemblies 2 and 4 of FIG. 1. Each of themagnet assemblies 4 and 2 includes three preferably concentric annularpermanent magnets 4a, 4b, 4c and 2a, 2b, 2c (not shown in drawing). Theannular permanent magnets 4a, 4b and 4c are offset from each other alongthe axis 12, and the annular permanent magnets 2a, 2b and 2c (not shown)are also offset from each other along the axis 12 as disclosed in U.S.Pat. No. 5,900,793 to Katznelson et al., now U.S. Pat. 5,900,793.

The axis 12 is the axis of symmetry of both magnet assemblies 2 and 4,passing through their centers. The axis 12 coincides with the z-axisalong which the main magnetic field generated by the magnet assemblies 2and 4 is oriented.

In order to reduce eddy currents each one of the concentric annularpermanent magnets 4a, 4b, 4c, 2a, 2b and 2c is formed from segments 24each of which is permanently magnetized in a known manner and thenattached to the neighboring segments using an electricallynon-conducting glue (not shown) or non-conductive spacers (not shown).For example, the segments 24 can be made from a neodymium-iron-boron(Nd--Fe--B) alloy. However, the segments 24 can be made from any otheralloy or ceramic material suitable for forming permanent magnets ofsufficient magnetic field strength. Preferably, the material from whichthe segments 24 are made should have a relatively low electricalconductivity.

The magnet assemblies 2 and 4 joined together by frame 3 (not shown inFIG. 2) define a region 16 having therein a volume 18 of substantiallyuniform magnetic field, between the pair of magnet assemblies 2 and 4.

The MRI probe 1 further includes Gradient coils (not shown) forgenerating gradient fields, shim coils (not shown) for active shimmingof the main magnetic field, RF coils (not shown), a temperature controlsystem (not shown) and an RF shield (not shown).

Ordinarily, the gradient fields are generated by a set of coils, throughwhich a current of an adequate magnitude flows. During the periods ofbuilding up and decay of the currents, the temporal change of themagnetic flux, originally generated by the currents, creates eddycurrents in conductive materials situated in their vicinity such as softiron parts or permanent magnet parts used in prior art MRI permanentmagnets or the aluminum enclosures of the cooling systems used insuper-conducting magnets of MRI systems. The eddy currents generated bythe gradient coil magnetic flux changes, generate secondary magneticfields which may interfere with the primary gradient fields and affecttheir precision in encoding the spatial information.

In prior art MRI devices, the gradient coils are located within theinternal free volume situated in the main magnet, where the imaged bodyis also introduced. To attenuate the effect of the spurious eddycurrents, prior art MRI devices may use shielded gradient coils orpre-emphasis circuits which modify gradient amplifier demand in order tocompensate for eddy currents. In small organ dedicated MRI probes and inMRI probes adapted for intra-operative use such as the MRI probe 1 ofFIG. 1, the dimensions of the region 16 (best seen in FIG. 2) foraccommodating the organ to be imaged are limited by practicalconsiderations. Generally, the design of such MRI systems involves atradeoff between maximizing the intensity and homogeneity of themagnetic field in as large an imaging volume as possible and providingmaximal accessibility of the surgeon to the organ undergoing surgery.For example, the MRI probe 1 (FIGS. 1 and 2) is designed to maximize thesize of the volume 18 of homogenous magnetic field while keeping thesize of the magnet assemblies 2 and 4 minimal while allowing enoughspace for positioning the shoulders of the patient 6. If one tries toincrease the space available for the shoulders of the patient 6 byincreasing the distance between the magnet assemblies 2 and 4 along theaxis 12, the resulting decrease in the strength and homogeneity of themagnetic field will have to be compensated. The magnetic field can becompensated by increasing the thickness of the annular permanent magnets4a, 4b, 4c of FIG. 2 and 2a, 2b and 2c (not shown in FIG. 2).

Increasing the thickness of the annular permanent magnets 4a, 4b, 4c,2a, 2b and 2c (not shown) is practically limited since their magneticfield depends non-linearly on their thickness. Thus, increasing thethickness of an annular permanent magnet above a certain value, resultsin a negligible contribution to the magnetic field strength.

The magnetic field can also be compensated by increasing the size anddiameter of the magnet assemblies 2 and 4. However, increasing thediameter of the magnet assemblies 2 and 4 may in turn shift the locationof the volume 18 relative to the desired position of the head of thepatient 6. The shifting may also prevent access to and imaging of thelower part of the brain, affecting the types of surgery that can beperformed using the probe 1.

Thus, placing the gradient coils and/or shim and RF coils within thealready restricted region 16 between the magnet assemblies 2 and 4,limits even further the space available for positioning the organ to beimaged and may hinder access to the organ undergoing surgery and theplacing and manipulating of surgical instruments within that organduring surgery.

Furthermore, in MRI systems using permanent magnets, if the gradientcoils are positioned in close proximity to the permanent magnets, theheat developed in the resistive gradient coils by the currents flowingwithin the coils may heat the permanent magnet. The heat generated bythe gradient coils may thus cause local temperature increase in thepermanent magnets. Such temperature changes are undesirable since thefield generated by permanent magnets is highly susceptible to largevariations induced by local temperature changes.

MRI systems based on permanent magnets such as the MRI probe 1 of FIG. 1or the MRI probe of FIG. 2, do not include electrically conductingstructures operating as magnetic flux return structures. This fact, inaddition to the segmented structure of the annular permanent magnets 4a,4b, 4c and 2a, 2b and 2c (not shown) and the intrinsic low conductivityof the Nd--Fe--B alloy from which they are made, substantially reducethe spurious eddy current problem.

Whole body MRI/MRT systems typically use a fixed installation RF cagefor preventing magnetic, electromagnetic and electrical noise from theoutside from penetrating into the imaging volume inside the probe andinterfering with the weak NMR signals generated during imaging. Inaddition, the RF cage is also used to reduce the leakage of the RFradiation generated within the probe during imaging to preventdisturbances to other electrical devices used near the MRI probe.

Unfortunately, for practical reasons, large fixed installation RF cagesor RF rooms cannot always be used small organ dedicated MRI or MRTprobes of the type used for intra-operative imaging such as the MRIprobe 1 of FIG. 1. For example, while the small organ dedicated MRIprobe 1 may be operated within a large shielded RF room, this willnecessitate the use of special expensive shielded surgical equipmentthat is designed to create minimal RFI disturbances so as not tointerfere with the operation of the MRI probe 1.

SUMMARY OF THE INVENTION

There is therefore provided, in accordance with a preferred embodimentof the present invention, electromagnetic apparatus for use in an MRIdevice. The probe includes a first permanent magnet assembly having afirst surface and a second surface thereof. The probe also includes asecond permanent magnet assembly having a third surface and a fourthsurface thereof. The second permanent magnet assembly opposes the firstpermanent magnet assembly such that the second surface and the thirdsurface define an open region therebetween, for producing apredetermined volume of substantially uniform magnetic field extendingin a first direction parallel to a first axis. The volume is disposedwithin the open region.

The probe also includes an energizable transmitting RF coil forproducing an RF electromagnetic field within the volume, an energizablez-gradient coil for producing a magnetic field gradient extending withinthe open region in the first direction and parallel to the first axis,an energizable x-gradient coil for producing a magnetic field gradientextending within the open region in parallel to a second axis orthogonalto the first axis, and an energizable y-gradient coil for producing amagnetic field gradient extending within the open region in parallel toa third axis orthogonal to the first axis and the second axis. At leastone of the x-gradient coil, y-gradient coil and z-gradient coil ispositioned outside of the open region.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the transmitting RF coil includes at least a firstportion thereof positioned within the open region adjacent the secondsurface and at least a second portion thereof positioned within the openregion adjacent the third surface. The first portion and the secondportion of the transmitting RF coil are electrically connected inseries.

Furthermore, in accordance with yet another preferred embodiment of thepresent invention, the transmitting RF coil further includes a thirdportion thereof including current return conductors positioned outsideof the open region and adjacent the first surface, and at least a fourthportion thereof including current return conductors positioned outsideof the open region and adjacent the fourth surface to increase theefficiency of the transmitting RF coil. The first portion, secondportion, third portion and fourth portion of the transmitting RF coilare electrically connected in series.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the apparatus further includes an energizable shimcoil for improving the homogeneity of the substantially uniform magneticfield.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the shim coil includes a first shim coil portionpositioned outside of the open region and opposed to the first surfaceof the first permanent magnet assembly, and a second shim coil portionpositioned outside of the open region and opposed to the fourth surfaceof the second permanent magnet assembly.

Further still, in accordance with another preferred embodiment of thepresent invention, the first shim coil portion and the second shim coilportion are electrically connected in series.

Furthermore, in accordance with another preferred embodiment of thepresent invention, at least one of the x-gradient coil, y-gradient coiland z-gradient coil includes a first coil portion thereof opposed to thefirst surface of the first permanent magnet assembly and a secondcomplementary coil portion thereof opposed to the fourth surface of thesecond permanent magnet assembly.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the first coil portion and the second coil portion ofthe at least one of the x-gradient coil, y-gradient coil and z-gradientcoil are electrically connected in series.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the first coil portion and the second coil portion ofat least one of the x-gradient coil, y-gradient coil and z-gradient coilare substantially planar printed circuits, the first coil portion isassembled into a first multi-layer printed circuit assembly opposed tothe first surface, and the second coil portion is assembled into asecond multi-layer printed circuit assembly opposed to the fourthsurface.

Furthermore, in accordance with another preferred embodiment of thepresent invention, each of the first multi-layer printed circuitassembly and second multi-layer printed circuit assembly furtherincludes a portion of an energizable shim coil, the portion of the shimcoil is a substantially planar printed circuit.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the apparatus further includes a mounting of lowpermeability material for mounting the first permanent magnet assemblyand the second permanent magnet assembly in opposition to each other.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the first permanent magnet assembly includes a firstannular permanent magnet with a first and a second surface thereof. Thefirst surface of the first annular permanent magnet is of a firstmagnetic polarity and the second surface of the first annular permanentmagnet is of a second magnetic polarity. The first annular permanentmagnet has an inside diameter. The first annular permanent magnet has atleast a portion of the first surface of the first annular magnet lyingin a first plane to provide a first magnetic field in the open region.The first magnetic field has a zero rate of change in a first directionat a first point in the open region. The first magnet assembly alsoincludes at least a second annular permanent magnet with a first and asecond surface thereof. The first surface of the second annular magnetis of the first magnetic polarity and the second surface of the secondannular permanent magnet is of the second magnetic polarity. The secondannular permanent magnet has an outside diameter which is smaller thanthe inside diameter of the first annular permanent magnet, with at leasta portion of the first surface of the second annular magnet lying in asecond plane spaced from the first plane to provide a second magneticfield whereby the second magnetic field is superimposed upon the firstmagnetic field in the open region, having a zero rate of change in thefirst direction at a second point different from the first point. Thesecond permanent magnet assembly includes a third annular permanentmagnet with a first and a second surface thereof, the first surface ofthe third annular permanent magnet is of the second magnetic polarityand the second surface of the third annular permanent magnet is of thefirst magnetic polarity. The third annular permanent magnet has aninside diameter, the third annular permanent magnet has at least aportion of the first surface of the third annular magnet lying in athird plane to provide a third magnetic field, whereby the thirdmagnetic field is superimposed on the first and second magnetic fieldsin the open region, having a zero rate of change in the first directionat a third point different from the first and second points. The secondmagnet assembly also includes at least a fourth annular permanent magnethaving a first and a second surface thereof, the first surface of thefourth annular magnet is of the second magnetic polarity and the secondsurface of the fourth annular permanent magnet is of the first magneticpolarity. The fourth annular permanent magnet has an outside diameterwhich is smaller than the inside diameter of the third annular permanentmagnet, with at least a portion of the first surface of the fourthannular permanent magnet lying in a fourth plane spaced from the thirdplane to provide a fourth magnetic field, whereby the fourth magneticfield is superimposed upon the first, second and third magnetic fields,in the open region, having a zero rate of change in the first directionat a fourth point different from the first, second and third points.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the first axis passes through the centers of thefirst annular permanent magnet, the at least second annular permanentmagnet, the third annular permanent magnet and the at least fourthannular permanent magnet.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the first annular permanent magnet, the at leastsecond annular permanent magnet, the third annular permanent magnet andthe at least fourth annular permanent magnet are rare-earth permanentmagnets.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the rare-earth permanent magnets areneodymium-iron-boron alloy permanent magnets.

Furthermore, in accordance with another preferred embodiment of thepresent invention, at least one of the first annular permanent magnet,the at least second annular permanent magnet, the third annularpermanent magnet and the at least fourth annular permanent includes aplurality of segments attached to adjacent segments using anelectrically non-conductive adhesive.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the segments are equiangular segments.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the segments have a trapezoidal cross-section in aplane orthogonal to the first direction.

Furthermore, in accordances with another preferred embodiment of thepresent invention, the z-gradient coil includes a first gradient coilportion concentrically disposed between the first annular permanentmagnet and the at least second annular permanent magnet, and a secondgradient coil portion concentrically disposed between the third annularpermanent magnet and the at least fourth annular permanent magnet. Thefirst and second gradient coil portions have their longitudinal axescoincident with the first axis.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the apparatus further including at least onereceiving RF coil placeable adjacent to an organ or body part disposedwithin the open region.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the transmitting RF coil is a linearly polarizing RFcoil.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the transmitting RF coil is a circularly polarizingRF coil.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the circularly polarizing RF coil is aquadrature-hybrid RF coil.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the first permanent magnet assembly includes a firstplurality of nested polygonally or elliptically shaped annular permanentmagnets, and the second permanent magnet assembly includes a secondplurality of nested polygonally or elliptically shaped annular permanentmagnets the first plurality being opposed to the second plurality suchthat the second plurality is configured as a mirror image of the firstplurality.

Furthermore, in accordance with another preferred embodiment of thepresent invention, at least one of the x-gradient coil, y-gradient coiland z-gradient coil is positioned below the first permanent magnetassembly and the second permanent magnet assembly.

Furthermore, in accordance with another preferred embodiment of thepresent intention, the x-gradient coil, the y-gradient coil and thez-gradient coil are planar printed circuit coil boards assembled withina single multi-layer printed circuit assembly positioned underneath thefirst permanent magnet assembly and the second permanent magnetassembly.

There is further provided, in accordance with a preferred embodiment ofthe present invention, electromagnetic apparatus for use in an MRIdevice. The apparatus includes a permanent magnet assembly having atleast a first surface defining a first side of the permanent magnetassembly and a second surface defining a second side of the permanentmagnet assembly opposed to the first side, for producing a predeterminedvolume of substantially uniform magnetic field extending in a firstdirection beyond the first surface. The apparatus further includes anenergizable transmitting RF coil for producing an RF electromagneticfield within the volume. At least a portion of the RF coil is positionedadjacent the first surface of the permanent magnet assembly. Theapparatus also includes an energizable z-gradient coil for producing amagnetic field gradient extending within the volume in the firstdirection parallel to a first axis. The apparatus also includes anenergizable x-gradient coil for producing a magnetic field gradientextending within the volume parallel to a second axis orthogonal to thefirst axis. The apparatus also includes an energizable y-gradient coilfor producing a magnetic field gradient extending within the volumeparallel to a third axis orthogonal to the first axis and to the secondaxis. At least one of the x-gradient coil, y-gradient coil andz-gradient coil is positioned opposing the second surface of thepermanent magnet assembly.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the apparatus further includes at least oneenergizable shim coil for improving the homogeneity of the substantiallyuniform magnetic field.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the at least one shim coil is a substantially planarcoil opposing the second surface of the permanent magnet assembly.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the x-gradient coil, the y-gradient coil and thez-gradient coil are substantially planar printed circuits assembledwithin a substantially planar multi-layer printed circuit assembly. Themulti-layer printed circuit assembly is disposed on the second side ofthe permanent magnet assembly facing the second surface.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the multi-layer printed circuit assembly furtherincludes at least one energizable shim coil. The at least one shim coilis a substantially planar printed circuit.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the permanent magnet assembly includes a firstannular permanent magnet having an upper and a lower surface thereof.The upper surface of the first annular permanent magnet is of a firstmagnetic polarity and the lower surface of the first annular permanentmagnet is of a second magnetic polarity. The first annular permanentmagnet has an inside diameter. The first permanent magnet has at least aportion of the upper surface of the first annular magnet lying in afirst plane and providing a first magnetic field in the predeterminedvolume. The first magnetic field has a zero rate of change in the firstdirection at a first point. The permanent magnet assembly furtherincludes at least a second annular permanent magnet having an upper anda lower surface thereof. The upper surface of the at least secondannular permanent magnet is of the first magnetic polarity and the lowersurface of the at least second annular permanent magnet is of the secondmagnetic polarity. The at least second annular permanent magnet has anoutside diameter which is smaller than the inside diameter of the firstannular permanent magnet. The at least second annular permanent magnetprovides a second magnetic field. The permanent magnet assembly alsoincludes low permeability material interconnecting the first annularpermanent magnet with the second annular permanent magnet, so that atleast a portion of the upper surface of the second annular permanentmagnet is in a second plane spaced from the first plane. The secondmagnetic field is superimposed upon the first magnetic field, in thepredetermined volume, having a zero rate of change in the firstdirection at a second point different from the first point.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the first axis passes through the center points ofthe first annular permanent magnet and the at least second annularpermanent magnet.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the first annular permanent magnet and the at leastsecond annular permanent magnet are rare-earth permanent magnets.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the rare-earth permanent magnets areneodymium-iron-boron alloy permanent magnets.

Furthermore, in accordance with another preferred embodiment of thepresent invention, at least one of the first annular permanent magnetand the at least second annular permanent magnet includes a plurality ofsegments attached to adjacent segments using an electricallynon-conductive adhesive.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the segments are equiangular segments.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the segments have a trapezoidal cross-section in aplane orthogonal to the first direction.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the z-gradient coil is an extended gradient coilconcentrically disposed between the first annular permanent magnet andthe at least second annular permanent magnet, the z-gradient coil has alongitudinal axis coincident with the first axis.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the apparatus further includes at least one receivingRF coil positioned on the first side of the permanent magnet assemblyand placeable adjacent to an organ or body part to be imaged using theapparatus.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the transmitting RF coil is a linearly polarizing RFcoil.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the transmitting RF coil is a circularly polarizingRF coil.

Furthermore, in accordance with another preferred embodiment of thepresent invention, at least a portion of the transmitting RF coil ispositioned on the second side of the permanent magnet assembly opposingthe second surface of the permanent magnet assembly to improve theefficiency of the transmitting RF coil.

There is also provided, in accordance with another preferred embodimentof the present invention, electromagnetic apparatus for use in an MRIdevice. The apparatus includes a permanent magnet assembly having afirst surface and a second surface for producing a predetermined volumehaving a magnetic field varying substantially linearly along a firstaxis. The volume extends in a first direction beyond the first surfacealong the first axis. The magnetic field is substantially uniform in anyplane which is included within the predetermined volume and which isorthogonal to the first direction within the predetermined volume. Theapparatus further includes an energizable transmitting RF coil fortransmitting RF radiation. The RF coil has at least one portion thereofpositioned opposing the first surface of the permanent magnet assembly.The apparatus also includes an energizable x-gradient coil for producinga magnetic field gradient along a second axis orthogonal to the firstaxis. The apparatus also includes an energizable y-gradient. coil forproducing a magnetic field gradient along a third axis orthogonal to thefirst axis and to the second axis. At least one of the x-gradient coiland y-gradient coil is positioned opposing the second surface of thepermanent magnet assembly.

Furthermore, in accordance with another preferred embodiment of thepresent invention, the apparatus further includes at least one receivingRF coil positioned on the first side of the permanent magnet assemblyand placeable adjacent to an organ or body part to be imaged using theapparatus.

There is also provided, in accordance with another preferred embodimentof the present invention, a method for constructing electromagneticapparatus for use in an MRI device. The method includes the steps ofproviding a first permanent magnet assembly having a first surface and asecond surface thereof, providing a second permanent magnet assemblyhaving a third surface and a fourth surface thereof, positioning thesecond permanent magnet assembly opposite the first permanent magnetassembly such that the second surface and the third surface define anopen region therebetween, for producing a predetermined volume ofsubstantially uniform magnetic field extending in a first directionparallel to a first axis, the volume is disposed within the open region,providing an energizable transmitting RF coil for producing an RFelectromagnetic field within the volume, providing an energizablez-gradient coil for producing a magnetic field gradient extending withinthe open region in the first direction and parallel to the first axis,providing an energizable x-gradient coil for producing a magnetic fieldgradient extending within the open region in parallel to a second axisorthogonal to the first axis, providing an energizable y-gradient coilfor producing a magnetic field gradient extending within the open regionin parallel to a third axis orthogonal to the first axis and the secondaxis, providing at least one receiving RF coil placeable adjacent to anorgan or body part to be imaged for receiving RF signals from the organor body part, and positioning at least one of the x-gradient coil,y-gradient coil and z-gradient coil outside of the open region forreducing the loading of the transmitting RF coil and the at least onereceiving RF coil by the at least one of the x-gradient coil, y-gradientcoil and z-gradient coil.

There is further provided, in accordance with another preferredembodiment of the present invention, a method for constructingelectromagnetic apparatus for use in an MRI device. The method includesthe steps of providing a permanent magnet assembly having at least afirst surface defining a first side of the permanent magnet assembly anda second surface defining a second side of the permanent magnet assemblyopposed to the first side, for producing a predetermined volume ofsubstantially uniform magnetic field extending in a first directionbeyond the first surface, providing an energizable transmitting RF coilfor producing an RF electromagnetic field within the volume, positioningat least a portion of the transmitting RF coil adjacent the firstsurface of the permanent magnet assembly, providing at least onereceiving RF coil placeable adjacent to an organ or body part to beimaged for receiving RF signals from the organ or body part, providingan energizable z-gradient coil for producing a magnetic field gradientextending within the volume in the first direction parallel to a firstaxis, providing an energizable x-gradient coil for producing a magneticfield gradient extending within the volume parallel to a second axisorthogonal to the first axis, providing an energizable y-gradient coilfor producing a magnetic field gradient extending within the volumeparallel to a third axis orthogonal to the first axis and to the secondaxis, and positioning at least one of the x-gradient coil, y-gradientcoil and z-gradient coil opposite the second surface of the permanentmagnet assembly for reducing the loading of the transmitting RF coil andthe at least one receiving RF coil by the at least one of the x-gradientcoil, y-gradient coil and z-gradient coil.

Finally, there is provided, in accordance with another preferredembodiment of the present invention, a method for constructingelectromagnetic apparatus for use in an MRI device. The method includesthe steps of providing a permanent magnet assembly having a firstsurface and a second surface for producing a predetermined volume havinga magnetic field varying substantially linearly along a first axis, thevolume extends in a first direction beyond the first surface along thefirst axis, the magnetic field is substantially uniform in any planeincluded within the predetermined volume and orthogonal to the firstdirection within the predetermined volume, providing an energizabletransmitting RF coil for transmitting RF radiation, positioning thetransmitting RF coil such that at least one portion thereof opposes thefirst surface of the permanent magnet assembly, providing at least onereceiving RF coil placeable adjacent to an organ or body part to beimaged for receiving RF signals from the organ or body part, providingan energizable x-gradient coil for producing a magnetic field gradientalong a second axis orthogonal to the first axis, providing anenergizable y-gradient coil for producing a magnetic field gradientalong a third axis orthogonal to the first axis and to the second axis,and positioning at least one of the x-gradient coil and y-gradient coilopposite the second surface of the permanent magnet assembly forreducing the loading of the transmitting RF coil and the at least onereceiving RF coil by the at least one of the x-gradient coil andy-gradient coil.

BRIEF DESCRIPTION OF THE DRAWINGS

The present invention will be described with reference to theaccompanying drawings, wherein like reference numerals identify like orcorresponding components.

In the drawings:

FIG. 1 is a schematic perspective view of a small organ dedicated MRIprobe useful in brain surgery;

FIG. 2 is a schematic isometric view illustrating the two permanentmagnet assemblies of FIG. 1;

FIG. 3 is a schematic cross section illustrating part of a prior art MRIdevice using permanent magnets;

FIG. 4 is an isometric view illustrating part of an MRI probe usingpermanent magnets and having external gradient coils, in accordance witha preferred embodiment of the present invention;

FIGS. 5-7 are front views schematically illustrating printed circuitlayout designs for an x-coil, y-coil and z-coil, respectively, useful inthe MRI probe of FIG. 4;

FIG. 8 is a schematic isometric view illustrating part of an MRI probehaving the z-gradient coil positioned in the volume between the twopermanent magnet assemblies and the x and y gradient coils positionedoutside of the volume between the two permanent magnet assemblies, inaccordance with yet another preferred embodiments of the presentinvention;

FIG. 9 is a cross section of the magnet assembly, printed circuit boardand multi-layer printed circuit assembly of FIG. 8 taken along the linesIX--IX;

FIG. 10 is a cross section illustrating part of an MRI probe havingz-gradient coils positioned between two annular permanent magnets, inaccordance with another preferred embodiment of the present invention;

FIG. 11 is an isometric view of part of the annular permanent magnet ofFIG. 2 useful in understanding the forces acting on a segment of theannular permanent magnet;

FIG. 12 is an isometric view illustrating part of an annular permanentmagnet including two layers of segments, in accordance with anotherpreferred embodiment of the present invention;

FIG. 13 is an isometric view illustrating part of an annular permanentmagnet including multiple layers having segments, in accordance with yetanother preferred embodiment of the present invention;

FIG. 14 is a pictorial illustration of a small organ dedicated MRI probeused in conjunction with a local disposable RF cage, in accordance witha preferred embodiment of the present invention;

FIG. 15 is a schematic isometric view illustrating a transmitting RFcoil providing linear polarization useful with the MRI probes of thepresent invention;

FIG. 16 is a schematic isometric view illustrating an MRI probeincluding the transmitting RF coil of FIG. 15 disposed therein, inaccordance with a preferred embodiment of the present invention;

FIG. 17 is a schematic cross section of the MRI probe of FIG. 16 takenalong the lines XVII--XVII;

FIG. 18 is a schematic cross section of an MRI probe having an internalZ-gradient coil and external X-gradient and Y-gradient coils, inaccordance with another preferred embodiment of the present invention;

FIG. 19 is a schematic cross section of an MRI probe having an internalZ-gradient coil and external X-gradient and Y-gradient coils, inaccordance with yet another preferred embodiment of the presentinvention;

FIG. 20 is a schematic isometric view illustrating an RF coil combinablewith the RF coil of FIG. 15 to form a circularly polarizing RFtransmitting coil assembly for use with an MRI probe, in accordance withanother embodiment of the present invention;

FIG. 21 is a schematic isometric view illustrating a circularlypolarizing RF transmitting coil assembly, assembled from the RF coil ofFIG. 15 and the RF coil of FIG. 21;

FIG. 22 is a schematic isometric view illustrating an MRI probe havingexternal X, Y and Z-gradient coils, in accordance with still anotherpreferred embodiment of the present invention;

FIG. 23 is a schematic diagram of an MRI probe having a single permanentmagnet assembly, in accordance with yet another preferred embodiment ofthe present invention; and

FIG. 24 is a schematic diagram illustrating an MRI probe having a fixedmagnetic field gradient, in accordance with another preferred embodimentof the present invention.

DETAILED DESCRIPTION OF THE DRAWINGS

Reference is now made to FIG. 3 which is a schematic cross sectionillustrating part of a prior art MRI device 30 using permanent magnets.The MRI device 30 includes two permanent magnets 36 and 38. Each of thepermanent magnets 36 and 38 is constructed from segments 34. Thepermanent magnets 36 and 34 are encased in a structure 32 made of aconducting metal such as soft iron and operating as a magnetic fluxreturn circuit. The MRI device 30 further includes two multi-layerprinted circuits 40 and 42 positioned in the volume between the twopermanent magnets 36 and 38. The multi-layer printed circuits includethe gradient coils. The MRI device 30 also includes RF coils (not shown)and the shim coils (not shown) of the MRI device 30. Each of themulti-layer printed circuits 40 and 42 is positioned in close proximityto the permanent magnets 36 and 38, respectively, such that enough roomis left for positioning the organ 44 such as the knee or head of apatient between the multi-layer printed circuits 40 and 42.

Reference is now made to FIG. 4 which is an isometric view illustratingpart of an MRI probe using permanent magnets and having externalgradient coils, in accordance with a preferred embodiment of the presentinvention.

The part of the MRI probe illustrated in FIG. 4 includes the two annularpermanent magnet assemblies 2 and 4 of FIG. 2 and two multi-layerprinted circuit assemblies 52 and 54. The multi-layer printed circuitassemblies 52 and 54 each include x, y and z-gradient coils (not shown),and shim coils (not shown). The MRI probe of FIG. 4 also includes RFcoils (not shown).

In contrast to the prior art Permanent magnet MRI device 30 of FIG. 3 inwhich the multi-layer printed circuits 40 and 42 including the gradientcoils are positioned in the volume between the two permanent magnets 36and 38, the multi-layer printed circuit assemblies 52 and 54 of FIG. 4are positioned outside of the region 14 defined between the twopermanent magnet assemblies 2 and 4.

It is noted that, the natural inclination of the designer is to placethe gradient coils of the multi-layer printed circuit assemblies 52 and54 between the magnet assemblies 2 and 4, placing them closer to theimaging volume. This will make the relative influence of the eddycurrents in the more distant magnet assemblies 2 and 4 on the gradientfields, smaller.

Additionally, in prior art MRI systems having a large structure of anelectrically conductive metal such as iron, which surrounds the magnetpoles, the gradient coils cannot be placed outside of the magnet polessince the conductive metal will absorb most of the gradient field.

Moreover, the gradient field strength is inversely proportional to thethird power of the distance of the gradient coil from the center of theimaging volume. If the permanent magnets are large, the designer willtend to place the gradient coils as close as possible to the imagingvolume, and avoid placing them outside the region between the permanentmagnets, since enormous amplifiers will be required to drive gradientcoils which are placed outside of the region between the permanentmagnets because of the low efficiency of the gradient coil.

However, if the multi-layer printed circuit assemblies 52 and 54 arepositioned between the magnet assemblies 2 and 4, they take up preciousspace between the magnet assemblies 2 and 4. Moreover, if themulti-layer printed circuit assemblies 52 and 54 are placed in closeproximity to the permanent magnet assemblies 2 and 4, respectively, theheat generated by the currents flowing in all the gradient coils withinthe multi-layer printed circuit assemblies 52 and 54 may raise thetemperature of the proximal surfaces of the magnet assemblies 2 and 4,affecting significantly the basic magnetic field.

The neodymium-iron-boron (Nd--Fe--B) alloy of which the segments 24 (notshown in FIG. 4) comprising the annular permanent magnets (not shown) ofthe permanent magnet assemblies 2 and 4 are made, features a relativelylow electrical conductivity. Hence, the eddy current effect isrelatively low, to begin with.

Moreover, the construction of the annular permanent magnets included inthe permanent magnet assemblies 2 and 4 from insulated segments 24 (bestseen in FIG. 2), reduces even more the influence of the eddy currenteffect because of the discontinuation of the current flow path in theannular permanent magnets. Thus, because the design of the magnetassemblies significantly reduces the eddy current spurious effects, itis possible to position the x, y and z gradient coils outside the region14, without significantly increasing the magnitude of the eddy currentseffects.

The design of the gradient coil printed circuit layout is well known inthe art and is not the subject of the present invention. Reference isnow briefly made to FIGS. 5, 6 and 7 which are front views schematicallyillustrating printed circuit layout designs for the x-gradient coil,y-gradient coil and z-gradient coil, respectively, useful inconstructing the multi-layer printed circuit assemblies 52 and 54 ofFIG. 4. It is noted that, the layout designs of FIGS. 5-7 are exemplarydesigns and that other designs can be used in implementing themulti-layer printed circuit assemblies of the present invention.

It in noted that, the printed circuits from which the multi-layerprinted circuit assemblies 52 and 54 are assembled are scaled in sizeand the magnitude of the currents is adapted in such a way, that thegradient fields of appropriate level are built up within the imagingvolume, without increasing the undesirable effects of the eddy currentson the main magnetic field and on the localized heating of the permanentmagnets 2 and 4.

An advantage of placing the multi-layer printed circuit assemblies 52and 54 outside the region 14, is the increase in available space in theregion 14 which is used for accommodating the imaged or treated organ.For example, the multi-layer printed circuit assemblies 52 and 54 aresufficiently distanced from the magnet assemblies 2 and 4, respectively,to allow free space for the shoulders (not shown in FIG. 4) of a patientundergoing brain surgery.

The freeing of extra space in the region 14 also results in increase inthe freedom and convenience of positioning and manipulating surgicalinstruments (not shown) or other equipment (not shown) within and aroundthe imaged or treated organ.

The eddy currents generated by the z-gradient coil, are the most likelyto interfere with the imaging process. The direction of the magneticfield generated by the z gradient coil is such that the eddy currentsresulting from changes in the magnetic flux generated by the z-gradientcreate a spurious magnetic field along the z-axis. Therefore, it may bedesirable to position the z-gradient coils in the region 14 between themagnet assemblies 2 and 4 while positioning the x-gradient coil, they-gradient coil and the shim coil outside of the region 14.

Reference is now male to FIG. 8 which is a schematic isometric viewillustrating part of an MRI probe having the z-gradient coil positionedin the volume between the two permanent magnet assemblies and having thex-gradient, y-gradient and the shim coils positioned outside of thevolume between the two permanent magnet assemblies, in accordance withyet another preferred embodiment of the present invention.

The part of the MRI probe illustrated in FIG. 8 includes the twopermanent magnet assemblies 2 and 4 of FIG. 2, a pair of multi-layerprinted circuit assemblies 72 and 74 positioned outside of the region 14and a pair of multi-layer printed circuit assemblies 76 and 78positioned in the region 14 between the magnet assemblies 2 and 4. Eachof the printed circuit boards 76 and 78 includes the z-gradient coils(not shown) as disclosed hereinabove. Each of the multi-layer printedcircuit assemblies 72 and 74 includes the x and y gradient coils and theshim coils (not shown).

Reference is now made to FIG. 9 which is a cross section of the magnetassembly 4 the printed circuit board 78 and the multi-layer printedcircuit assembly 74 of FIG. 8 taken along the lines X--X. The printedcircuit board 78 is positioned within the recess 75 formed by theannular permanent magnets 4a and 4b which are offset along the axis 12.The printed circuit board 76 (best seen in FIG. 8) is similarlypositioned within a recess (not shown) formed by the annular permanentmagnets 2a and 2b (FIG. 8) which are also offset along the axis 12.

Thus, the printed circuit boards 76 and 78 do not occupy a significantportion of the space in the region 14 between the magnet assemblies 2and 4.

The arrangement of the x, y and z gradient coils illustrated in FIGS. 8and 9 has the advantage that the z-gradient is generated close to theimaging volume 18 (best seen in FIGS. 2 and 9). Thus, the superimposedz-gradient field is minimally affected by the spurious eddy currents,while keeping the region 14 minimally obstructed by positioning the xand y-gradient coils outside of the region 14.

Reference is now made to FIG. 10 which is a cross section illustratingpart of an MRI probe having z-gradient coils positioned between twoannular permanent magnets, in accordance with another preferredembodiment of the present invention.

The magnet assembly 63 includes the annular permanent magnets 63a, 63band 63c. In contrast to the MRI probe of FIG. 8 in which the z-gradientcoils are substantially planar coils included in the multi-layer printedcircuit assemblies 72 and 74, the z-gradient coil 65 of FIG. 10 ispositioned between the annular permanent magnets 63a and 63b and extendsin a direction parallel to the axis 12. The multi-layer printed circuitassembly 64 includes x-gradient and y-gradient coils (not shown) and theshim coil (not shown). The multi-layer printed circuit assembly 62 mayalso include the RF coils (not shown).

It is noted that, the z-gradient coil 65 may be shaped as a helical coilwith a circular cross section or as an extended coil having anothersuitable cross section such as a regular polygonal cross section.

It is further noted that, while the z-gradient coil 65 of FIG. 10 isshown as an even-pitched helical coil for the sake of clarity ofillustration, the z-gradient coil 65 may have a variable pitch such thatthe number of coil windings per unit length may vary along the coil asis well known in the art.

It will be appreciated by those skilled in the art, that many variationshaving different positioning of the gradient coils and the shim coilsare possible which are within the scope of the present invention. Forexample, in all the preferred embodiments of FIGS. 4, 8, 9 and 10, theshim coils may be positioned in the region 14 between the magnetassemblies 2 and 4 (FIGS. 4, 8 and 9), or between the magnet assembly 63(FIG. 10) and the complementary magnet assembly (not shown in FIG. 10).

It is noted that, in all the embodiments of the present invention, theRF coils are positioned in the region 14 between the pairs of permanentmagnet assemblies. However, in other preferred embodiments of thepresent invention, the receiving RF coil (not shown) can be internallypositioned within the region 14 while a portion of the transmitting RFcoil (not shown) is positioned outside the region 14.

Returning to FIG. 3, the structure 32 serving as the magnetic fluxreturning circuit encases the segments 34 of the permanent magnets 36and 38. If the glue between any of the segments 34 becomes loose, theattractive force between the segments 34 of the permanent magnets 36 and38 will tend to pull the loose segments towards each other and into theregion where the organ of the patient 44 is positioned. However, thestructure 32 encases the segments 34 and prevents any loose segments 34from moving towards the organ of the patient 44.

Turning briefly to FIGS. 2 and 4, If any of the segments 24 of themagnet assemblies 2 and 4 become loose or fractured, the loose segmentsor loose fragments thereof may fly into the region 16 (FIG. 2) or region14 (FIG. 4) because of the attractive forces between the segments of themagnet assemblies 2 and 4.

Reference is now made to FIG. 11 which is an isometric view of part ofThe annular permanent magnet 4a of FIG. 2 illustrating the forces actingon a segment of the annular permanent magnet 4a.

Each of the segments 24 of the permanent annular magnet 4a has a planarsurface 24A facing towards the magnet assembly 2 (not shown), a planarsurface 24B parallel to the surface 24A and facing away from the magnetassembly 2, and a curved surface 24C. Each of the segments 24 of thepermanent annular magnet 4a is subjected to a force F1 which is normalto the surface 24A. The force F1 is due to magnetic attraction betweenthe segment 24 and the magnet assembly 2 (FIG. 2). Each of the segments24 of the permanent annular magnet 4a is also subjected to a force F2which is normal to the plane tangential to the surface 24C. The force F2is the vector sum of the repulsive forces (not shown) between thesegment 24 and all the other segments 24 of the annular permanent magnet4a. These repulsive forces arise due to the fact that all of thesegments 24 included within the annular permanent magnet 4a have theirmagnetic axes aligned parallel to each other. The direction and polarityof the magnetic field of each of the segments 24 is indicated by thestippled arrows labeled S and N indicating the south and north poles,respectively, of each of the segments 24.

If the glue attaching any of the segments 24 fails or if any of thesegments fractures, the segments or the fragments thereof may fly in adirection of the vector sum of the forces F1, F2 and any additionalforces (not shown) acting on the segments or on the fragments thereof.

Reference is now made to FIG. 12 which is an isometric view illustratingpart of an annular permanent magnet including two layers of segments, inaccordance with another preferred embodiment of the present invention.

The annular permanent magnet 84 includes two layers 86 and 88. The layer86 includes a plurality of segments labeled 87 and the layer 88 includesa plurality of segments labeled 89. The segments 87 of the layer 86 areattached to each other by an electrically non-conducting glue (notshown) or by electrically non-conducting spacers (not shown). Thesegments 89 of the layer 88 are also attached to each other by anelectrically non-conducting glue (not shown) or by electricallynon-conducting spacers (not shown). The layers 86 and 88 are attached toeach other by an electrically non-conducting glue (not shown) or by a anelectrically non-conducting spacer or spacers (not shown). Preferably,the shape and dimensions of the layers 86 and 88 are identical and theshape and dimensions of each of the segments 87 and 89 are identical.The layers 86 and 88 are attached to each other such that the segments87 are symmetrically staggered with respect to the segments 89. Thealignment of the layer 86 relative to layer 88 is such that each of thesegments 87 is offset from the underlying segment 89 by a distanceequivalent to half the arc subtended by the segment 87 along thecircumference of the circle 100. The direction and polarity of themagnetic field of each of the segments 87 and 89 is indicated by thestippled arrows labeled S and N as disclosed hereinabove. The annularpermanent magnet 84 is part of a magnet assembly (not shown) which isaligned with another similarly constructed magnet assembly (not shownfor the sake of clarity of illustration), this alignment is similar tothe alignment of the magnets assemblies 2 and 4 of FIGS. 2 and 4 withrespect to the alignment axis and the polarity of the magnetic fields ofthe magnet assemblies 2 and 4.

Each of the segments 87 is subjected to a force F3 which is normal tothe surface 87a. The force F3 is due to the magnetic attraction betweenthe segment 87 and the magnet assembly (not shown) aligned opposite tothe magnet assembly including the annular permanent magnet 84. Each ofthe segments 87 of the permanent annular magnet 84 is also subjected toa force F4 which is normal to the plane tangential to the surface 87c.The force F4 is the vector sum of the repulsive forces (not shown)between the segment 87 and all the other segments 87 of the annularpermanent magnet 84. These repulsive forces arise due to the fact thatall of the segments 87 included within the annular permanent magnet 84act as magnetic dipoles aligned parallel to each other and having thesame south-north orientation.

However, in contrast to the segments 24 of FIG. 11, each of the segments87 is subjected to a force F5 which is normal to the surface 87a andopposite in direction to the force F3. The force F5 is due to themagnetic attraction between the segment 87 and the underlying segments89 of the layer 88.

The force F5 is much larger than the vector sum of the forces F3 and F4.

Each of the segments 89 is subjected to a set of forces similar to theset of forces acting on the segments 87, except that the magnitude ofthe force (not shown) attracting the segment 89 towards the opposingmagnet assembly (not shown) is smaller then F3 due to the greaterdistance between the segments 89 and the opposing magnet assembly, andthe force (not shown) attracting the segment 89 to the segments 87 isequal in magnitude but opposite in direction to the force F5.

Thus, if the glue holding any of the segments 87 and 89 loosens or anyof the segments fracture, the two layers 86 and 88 including all thesegments or possible fragments thereof are attracted to each other by anet attractive force which will prevent the loose segments or fragmentsfrom flying or falling. This net attractive force holds any loosesegment or fragment attached in its place.

An advantage of the staggered double layer structure of the annularpermanent magnet of FIG. 12 is the increased safety for the patientwhose organ is imaged or treated in the MRI probe of the presentinvention.

An additional benefit of the staggered double layer structure of FIG. 12is the improved homogeneity of the magnetic field generated by theannular permanent magnet. This improved homogeneity results from thefact that the thin layers (not shown) of electrically non-conductingglue or spacer between the segments 24 of FIG. 11 cause local reductionsin the magnetic field intensity since they are made from non-magneticmaterial. The staggered double layer structure of the annular permanentmagnet 84 illustrated in FIG. 12 reduces the magnitude of these localreductions by approximately 50% relative to the local reductions of thesingle layer annular permanent magnet 4a of FIG. 11, for annularpermanent magnets 4a and 84 having the same dimensions.

This reduction occurs because the height of the non-magnetic glue orspacer material present within the annular permanent magnet 84 along thedirection of the magnetic field is reduced by half because of thestaggering of the segments 87 and 89 disclosed hereinabove.

It is noted that, the layer arrangement within the annular permanentmagnets of the present invention is not limited to two layers and thatmultiple layer arrangements are also possible.

Reference is now made to FIG. 13 which is an isometric view illustratingpart of an annular permanent magnet including multiple layers ofsegments, in accordance with yet another preferred embodiment of thepresent invention.

The annular permanent magnet 104 includes three layers 106, 108 and 110.Preferably, the layers 106, 108 and 110 are identical in size and shapeand include segments 107 which are preferably identical in size. Thesegments 107 of each of the layers 106, 108 and 110 are attached to eachother with electrically non-conducting glue or electricallynon-conducting spacers (not Shown) and the layers 106, 108 and 110 arealso attached to each other with electrically non-conducting glue orelectrically non-conducting spacers (not Shown) as disclosedhereinabove. The segments 107 of the layer 106 are staggered withrespect to the segments 107 of the layer 108 such that each of thesegments 107 of the layer 106 is offset from the underlying segment 107of the layer 108 by a distance equivalent to a third of the arcsubtended by the segment 107 along the circumference of the circle 109.The segments 107 of the layer 108 are staggered with respect to thesegments 107 of the underlying layer 110 such that each of the segments107 of the layer 108 is offset from the underlying segment 107 of thelayer 110 by a distance equivalent to a third of the arc subtended bythe segment 107 along the circumference of the circle 109.

The direction and polarity of the magnetic field of each of the segments107 is indicated by the stippled arrows labeled S and N as disclosedhereinabove.

This arrangement, even further reduces the inhomogeneity of the magneticfield generated by the annular permanent magnet 104, while still havinga net force attracting the segments 107 to each other as disclosedhereinabove.

It is noted that, the annular permanent magnet of the present inventionmay be constructed from another number of staggered layer of segmentsgreater than three. However, there may be a practical limit to thenumber of layers implemented due to the increasing difficulty ofassembling, aligning and gluing of a large number of segments.

It is further noted that while the preferred embodiments of FIGS. 12 and13 illustrate annular permanent magnets having layers of identicalheight along the direction of the magnetic field, other implementationsare possible using various combinations of layers having differentheights.

It is still further noted that, while the preferred embodiments of FIGS.12 and 13 show the detailed structure of only one of the annularpermanent magnets 84 and 104 which are part of magnet assemblies (notshown) containing multiple concentric annular permanent magnets, theother annular permanent magnets included in the magnet assemblies arepreferably also structured from multiple staggered layers as disclosedhereinabove. Moreover, the number of layers and the type of staggeringfor each of the annular permanent magnets composing a single magnetassembly may be selected according to the desired magnetic fieldproperties and manufacturing considerations.

U.S. Pat. No. 5,900,793 to Katznelson et al. disclosed hereinaboveteaches a method of improving the homogeneity of the magnetic fieldbetween opposing annular permanent magnets used in an MRI probe. Themethod includes selecting segments from a batch of equi-angular segmentsso that the variations in a magnetic field strength of adjacent segmentsfollow a cyclic curve having a regular period, and combining thesegments to form an annular permanent magnet. A magnet assembly isformed by connecting two or more such annular permanent magnets by a lowmagnetic permeability material. Finally two such magnet assemblies arealigned such that for each pair of opposing annular permanent magnets,the cyclic curves representing the magnetic field variation are alignedin anti-phase. The method improves the homogeneity of the achievablepermanent magnetic field.

In accordance with yet another preferred embodiment of the presentinvention, the method disclosed by Katznelson et al. in U.S. Pat. No.5,900,793 can be similarly applied in constructing annular permanentmagnets used in the MRI probes of the present invention.

Reference is now made to FIG. 14 which is a pictorial illustration of asmall organ dedicated MRI probe used in conjunction with a localdisposable RF cage, in accordance with a preferred embodiment of thepresent invention.

The small organ dedicated MRI probe 120 includes the magnet assemblies 2and 4 of FIG. 1 which are attached to a surgical table 122 by anadjustable frame 123. The surgical table is made from a conductivematerial such as stainless steel. The probe 120 also include an RF cage124. The RF cage 124 is made of a sheet of flexible conductive RF meshhaving a size and shape suitable for covering the body of the patient 6and the magnet assemblies 2 and 4.

The RF cage 124 is electrically connected to the surgical table 122 forcompleting the shielding of the MRI probe 120. The RF cage 124 may besuitably grounded. The RF cage 124 also has an opening 126 therein. Theopening 126 is used by the surgeons 130 and 132 for accessing the brainof the patient 6 during surgery. For example, a surgical instrument 128can be inserted through the opening 126 into the brain of the patient 6.The size and shape of the opening 126 is designed to enable comfortableinsertion and manipulation of surgical instruments therethrough, whilestill enabling effective shielding of the MRI probe creating a "selfshielded" magnet.

The RF cage 124 can also be made from a conductive flexible RF mesh madeof a conductive material such as copper metal embedded in a flexiblesheet of art electrically non-conductive material such as a suitableplastic. The RF cage 124 can also be made from a thin flexible sheet ofconductive material such as metal foil.

It is noted that, the materials from which the RF cage is made must besterilizable and the RF cage 124 is disposable.

It is further noted that, the method of placing some or all of thegradient coils outside of the region containing the imaging volume mayenable further increasing of the available space between the permanentmagnet assemblies of the present invention by properly configuring andpositioning of the RF coil.

Reference is now made to FIG. 15 which is a schematic isometric viewillustrating a transmitting RF coil providing linear polarization usefulwith the MRI probes of the present invention.

The transmitting RF coil 140 is preferably made of a folded flat copperribbon conductor but can be made of any other suitably shapedelectrically conducting material capable of carrying the requiredelectrical currents. The coil 140 includes four front conductor portions142A, 142B, 144A and 144B. When the RF coil 140 is electricallyenergized, an electrical current flows therethrough in the directionindicated by the arrows. The four front conductors 142A, 142B, 144A and144B effectively form an open Helmholz coil configuration suitable forgenerating a linearly polarized RF electromagnetic field.

The coil 140 also includes four current return conductor portions 152A,152B, 154A and 154B.

The coil terminals 148A and 148B are electrically connected to asuitable RF amplifier (not shown) for energizing the coil 140.

Reference is now made to FIGS. 16 and 17. FIG. 16 is a schematicisometric view illustrating an MRI probe 150 including the transmittingRF coil 140 of FIG. 15 disposed therein, in accordance with a preferredembodiment of the present invention. FIG. 17 is a schematic crosssection of the MRI probe 150 of FIG. 16 taken along the linesXVII--XVII.

The MRI probe 150 includes two permanent magnet assemblies 162 and 164.The permanent magnet assembly 162 includes a housing 182 and a set ofthree concentric annular permanent magnets 2A, 2B and 2C attached to thehousing 182. The permanent magnet assembly 164 includes a housing 184and a set of concentric annular permanent magnets 4A, 4B and 4C attachedto the housing 184. The housings 182 and 184 are made of fiberglass orfrom any other suitable electrically non-conducting plastic material orthe like. The details of the structure, construction and tuning of theannular permanent magnets included within the permanent magnetassemblies 162 and 164 are not the subject matter of the presentinvention and will therefore not be discussed in detail herein. Thestructure and design of such permanent magnet assemblies is disclosed inco-pending U.S. Pat. No. 5,900,793 to Katznelson et al.

The permanent magnet assembly 162 includes a first surface 182A and asecond surface 182B. The permanent magnet assembly 164 includes a thirdsurface 184A and a fourth surface 184B. The two permanent magnetassemblies 162 and 164 are attached to a frame 173 and oppose each othersuch that the second surface 182B and the third surface 184A definetherebetween an open region of space 114. The permanent magnetassemblies 162 and 164 produce a region of substantially homogenousmagnetic field 168 disposed within the region 114.

The probe 150 also includes multi-layer printed circuit assemblies 172and 174. The printed circuit assemblies 172 and 174 each include planarprinted circuits boards (not shown) comprising an X-gradient coil, aY-gradient coil, a Z-gradient coil and shim coils as disclosed in detailhereinabove for the multi-layer printed circuit assemblies 52 and 54 theMRI probe of FIG. 4. The printed circuit assembly 172 is disposedoutside of the region 114 and faces the first surface 182A of thepermanent magnet assembly 162. The printed circuit assembly 174 is alsodisposed outside of the region 114 and faces the fourth surface 184B ofthe permanent magnet assembly 164.

The MRI probe 150 further includes a transmitting RF coil 140 forproducing an RF electromagnetic field within the open region 114, and areceiving RF coil 175 positioned within the open region 114, adjacent tothe organ or body part (not Shown) which is to be imaged, for receivingRF electromagnetic signals from the organ or body part.

It is noted that the receiving RF coil 175 can be any suitable type ofreceiving RF coil known in the art, such as a flexible RF coil (notshown) or other types of RF coils having suitable dimensions forpositioning near the organ or body part which is being imaged.Furthermore, in accordance with another embodiment of the presentinvention the MRI probe 150 may also include a plurality of smallreceiving RF coils (not shown) which may be disposed at differentpositions adjacent the organ or body part (not shown) as is well knownin the art.

The part of the transmitting RF coil 140 which includes the four frontconductor portions 142A, 142B, 144A and 144B is disposed in the openregion 114 between the permanent magnet assemblies 162 and 164.

The front conductor portions 144A and 144B are positioned adjacent tothe surface 182B and are glued or attached thereto. Similarly, the frontconductor portions 142A and 142B are positioned adjacent to the surface184A and are glued or attached thereto. However, the front conductorportions 144A, 144B and 142A, 142B may also be positioned adjacent tothe surfaces 182B and 184A, respectively, without being attachedthereto.

Preferably, in accordance with a design for an open Helmholz coil thedistance between the front conductor portions 144A and 144B and thedistance between the front conductor portions 142A and 142B is designedsuch that α=60°, wherein α is the angle between the lines connecting thecenter point 169 of the imaging volume 168 with the centers of the frontconductor portions 142A and 142B. The point 169 lies on the axis 12 andis the midpoint between the surfaces 182B and 184A. However, the angle αmay also be different than 60° depending, inter alia, on the particulardesign parameters of the transmitting RF coil.

The part of the transmitting RF coil 140 which includes the four currentreturn conductor portions 152A, 152B, 154A and 154B is disposed outsideof the open region 114. The current return conductor portions 152A and152B are disposed between the surface 184B and the multi-layer printedcircuit assembly 174, and the current return conductor portions 154A and154B are disposed between the surface 182A, and the multi-layer printedcircuit assembly 172.

An advantage of the above design of the transmitting RF coil 140 is thatThe current return conductor portions 152A, 152B, 154A and 154B which donot contribute to the RF field and may actually cause a reductionthereof can be distanced from the corresponding front conductor portions142A, 142B, 144A and 144B to diminish the RF field reduction by thecurrent return conductor portions 152A, 152B, 154A and 154B.

An additional advantage of disposing the current return conductorportions 152A, 152B, 154A and 154B outside the region 114 is theincrease in the space available within the open region 114 forpositioning and manipulating an organ to be imaged or surgicalinstruments during medical interventional procedures.

It is noted that, the positioning of the multi-layer printed circuitassemblies 174 and 172 outside the region 114 and away from the frontconductor portions 142A, 142B 144A and 144B, significantly reduces theloading of the transmitting RF coil 140 by the X, Y and Z coils (notshown) and the shim coils (not shown) which are disposed within themulti-layer printed circuit assemblies 174 and 172. The annularpermanent magnets 2A, 2B, 2C, 4A, 4B and 4C have a lower electricalconductivity than the copper conductors of the X,Y,Z coils and the shimcoils, because they are made of a material, such as aniron-neodimium-boron alloy, having electrical conductivity lower thancopper and because of the construction of each of the annular permanentmagnets 2A, 2B, 2C, 4A, 4B and 4C from a plurality segments which areelectrically isolated from each other by an electrically non-conductingglue as disclosed in detail in co-pending U.S. Pat. No. 5,900,793 toKatznelson et al. Thus, the loading of the transmitting RF coil 140 andof the receiving RF coil 175 is significantly reduced by the placementof the multi-layer printed circuit assemblies 174 and 172 outside theregion 114 and away from transmitting RF coil 140 and the receiving RFcoil 175. The reduction in loading of the transmitting RF coil 140enables achieving a desired transmitted signal quality without having touse expensive high-power RF transmitting Amplifiers. The reduction inloading of the receiving RF coil 175 enables achieving a significantimprovement in the image quality obtained by the MRI probe 150.

It will be appreciated by those skilled in the art that, although thecurrent return conductor portions 152A, 152B and 154A, 154B are beingpositioned closer to the multi-layer printed circuit assemblies 174 and172, respectively, by being disposed outside of the open region 114,thus, potentially increasing the loading of the RF coil 140 by thegradient coils and shim coils, the multi-layer printed circuitassemblies 174 and 172 can be sufficiently distanced from the currentreturn conductor portions 152A, 152B and 154A, 154B, respectively, bymoving the multi-layer printed circuit assemblies 174 and 172 along theaxis 12 away from the point 169 to reduce the loading of the RF coil140.

The design of the MRI probe 150 can be thus optimized to give a desiredhigh image quality by reducing the loading of the RF coil 140 withouthaving to excessively increase the distance of the multi-layer printedcircuit assemblies 174 and 172 from the point 169 of the imaging volume168 which will require the use of stronger and more expensive amplifiersto energize the gradient and shim coils.

It is noted that, in prior art large MRI devices such as whole bodyimaging MRI devices, the gradient coils and the transmitting RF coilsare internally disposed in the region between the magnets. Typically,this region is large enough to allow designing a sufficient distancebetween the transmitting RF coil and the gradient coils, thus solvingthe problem of reducing the loading the RF coil by the gradient and/orshim coil.

In direct contrast, in the smaller and more compact MRI probes used insystems such as the interventional MRI/MRT systems of the presentinvention, the problem of loading of the transmitting RF coil is moredifficult to solve because the region between the permanents magnets(such as the regions 14 and 114 of FIGS. 4 and 17, respectively) issmall due to limitations on the allowable size of the permanent magnetassemblies. Thus, the use of external gradient and shim coils of thepresent invention which are placed outside the region between the magnetassemblies, has the advantage of making more space available between thepermanent magnet assemblies as well as reducing the loading of thetransmitting RF coil for improving the image quality attainable.

It is noted that, while in the MRI probe 150 of FIGS. 16 and 17 theX-gradient coil, Y-gradient coil, Z-gradient coil and shim coils areincluded within the multi-layer printed circuit assemblies 174 and 172which are externally positioned outside the region 114, other preferredembodiments of the present invention are possible in which some of thegradient coils and/or the shim coils are internally positioned withinthe region between the permanent magnet assemblies 164 and 162.

Reference is now made to FIG. 18 which is a schematic cross section ofan MRI probe 250 having an internal Z-gradient coil and externalX-gradient and Y-gradient coils, in accordance with another preferredembodiment of the present invention. The MRI probe 250 includes twoexternal multi-layer printed circuit assemblies 274 and 272 and twopermanent magnet assemblies 262 and 264. The multi-layer printed circuitassemblies 274 and 272 are similar in construction to the multi-layerprinted circuit assemblies 174 and 172 of FIG. 17, except that they donot include a Z-gradient coil. Thus, each of the multi-layer printedcircuit assemblies 274 and 272 includes an X-gradient coil (not shown),a Y-gradient coil (not shown) and a shim coil (not shown).

The two permanent magnet assemblies 262 and 264 are similar to the twopermanent magnet assemblies 162 and 164 of FIG. 17, except thatpermanent magnet assembly 262 also includes a printed circuit board 200which includes a Z-gradient coil (not shown) disposed in the spacebetween the housing 182 and the annular permanent magnets 2B and 2C, andpermanent magnet assembly 264 also includes a printed circuit board 202which includes a Z-gradient coil (not shown) disposed in the spacebetween the housing 182 and the annular permanent magnets 4B and 4C.

The placement of the printed circuit boards 200 and 202 inside thehousings 182 and 184, respectively, does not diminish the spaceavailable between the surfaces 182B and 184A, leaving the region 114free for positioning an organ or body part for imaging and enablingaccess of surgical instruments to the organ or body part.

Reference is now made to FIG. 19 which is a schematic cross section ofan MRI probe 350 having an internal Z-gradient coil and externalX-gradient and Y-gradient coils, in accordance with yet anotherpreferred embodiment of the present invention.

The MRI probe 350 includes the multi-layer printed circuit assemblies274 and 272 of FIG. 18 and permanent magnet assemblies 362 and 364. Thepermanent magnet assemblies 362 and 364 are identical to the permanentmagnet assemblies 162 and 164 of FIG. 17 in all respects except that thepermanent magnet assembly 362 also includes an extended Z-gradient coil300 and that that the permanent magnet assembly 364 also includes anextended Z-gradient coil 302. The Z-gradient coil 300 is concentricallydisposed between the annular permanent magnets 2A and 2B and theZ-gradient coil 302 is concentrically disposed between the annularpermanent magnets 4A and 4B.

The placement of the Z-gradient coils 300 and 302 inside the housings182 and 184, respectively, does not diminish the space available betweenthe surfaces 182B and 184A, leaving the region 114 free for positioningan organ or body part for imaging and enabling access of surgicalinstruments to the organ or body part.

It is noted that, the z-gradient coils 300 and 302 may be shaped as ahelical coil with a circular cross section or as an extended coil havinganother suitable cross section such as a regular polygonal cross section(not shown). The pitch of the coil windings may vary in accordance withthe required gradient parameters.

It is further noted that, while the z-gradient coils 300 and 302 of FIG.19 are shown as even-pitched helical coils for the sake of clarity ofillustration, the z-gradient coils 300 and 302 may have a variable pitchsuch that the number of coil windings per unit length may vary along thecoil as is well known in the art.

It is further yet noted that, the Z-gradient coil 65 of FIG. 10 and theZ-gradient coils 300 and 302 of FIG. 19 may also be shaped as extendedcoils having variable diameter windings, such that some portions of thecoil may have a different diameter than other portions of the coil.

Furthermore, it is noted that, for the sake of clarity of illustration,the receiving RF coil 175 of FIG. 16 is not shown in FIGS. 17-19

It is still further noted that, while the transmitting RF coil 140 ofthe MRI probe 150 is a linearly polarizing, other types of transmittingRF coils may be used.

Reference is now made to FIGS. 20 and 21. FIG. 20 is a schematicisometric view illustrating an RF coil 240 combinable with the RF coil140 of FIG. 15 to form a circularly polarizing RF transmitting coilassembly for use with an MRI probe, in accordance with anotherembodiment of the present invention. FIG. 21 is a schematic isometricview illustrating a circularly polarizing RF transmitting coil assembly,assembled from the RF coil of FIG. 15 and the RF coil of FIG. 20.

The RF coil 240 of FIG. 20 is preferably made of a folded flat copperribbon conductor but can be made of any other suitably shapedelectrically conducting material capable of carrying the requiredelectrical currents. The coil 240 includes four front conductor portions242A, 242B, 244A and 244B. When the RF coil 240 is electricallyenergized, an electrical current flows therethrough in the directionindicated by the arrows. The four front conductors 242A, 242B, 244A and244B effectively form an open Helmholz coil configuration.

The coil 240 also includes four current return conductor portions 252A,252B, 254A and 254B. It is noted that, while the front conductorportions 142A, 142B, 144A and 144B and the current return conductorportions 152A, 152B, 154A and 154B of the transmitting RF coil 140 (FIG.15) are aligned vertically, the front conductor portions 242A, 242B,244A and 244B and the current return conductor portions 252A, 252B, 254Aand 254B of the RF coil 240 are horizontally aligned.

The coil terminals 248A and 248B are electrically connected to asuitable RF amplifier (not shown) for energizing the coil 240.

The transmitting RF coil 240 of FIG. 20 may replace the transmitting RFcoil 140 of the MRI probe 150 (FIG. 16). However, the transmitting RFcoils 140 and 240 can also be combined to form the circularly polarizingtransmitting RF coil 340 of FIG. 21.

In the transmitting RF coil 340, the front conductor portions 242A and242B are aligned orthogonal to the front conductor portions 142A and142B. The four front conductor portions 242A, 242B, 142A and 142B aredisposed adjacent to the surface 184A of the permanent magnet assembly164 (not shown) of the MRI probe. The front conductor portions 244A and244B are aligned orthogonal to the front conductor portions 144A and144B. The four front conductor portions 244A, 244B, 144A and 144B aredisposed adjacent to the surface 182A of the permanent magnet assembly162 (not shown) of the MRI probe.

Care is taken to prevent any electrical contact between any portions ofthe RF coil 140 and portions of the RF coil 240, in order to preventshort circuits. This may be done by isolating the surface of the RFcoils 140 and 240 by a layer or coat of an insulating material (notshown) or by separating regions of possible contact with pieces ofelectrically non-conducting material.

The general design of circularly polarizing transmitting RF coils isknown in the art as a quadrature-hybrid RF coil type. However, theinventors of the present invention have noted that by positioning someor all of the gradient coils and shim coils outside of the open region114 the load on the transmitting RF coils can be significantly reducedand the RF coil efficiency is improved. Additionally, the positioning ofthe current return conductor portions 154A, 154B, 254A, 254B, 152A,152B, 252A and 252B of the circularly polarizing transmitting RF coil340 outside of the open region 114 additionally improves the RF coilefficiency by significantly increasing the distance of the currentreturn conductor portions 154A, 154B, 254A, 254B, 152A, 152B, 252A and252B from the open region 114.

It is noted that, while the transmitting RF coils 140 and 340 of FIGS.15, 16 and 21 which are useful with the MRI probes of the presentinvention have the advantage that portions thereof such as the currentreturn conductor portions are disposed outside the open region 114 toincrease the space available therewithin, many other designs of linearlyor circularly polarizing transmitting RF coils may be possibly used withMRI probes having external gradient coils disposed outside of the openregion 114, which are within the scope and spirit of the presentinvention. For example, transmitting RF coils (not shown) in which allof the transmitting RF coil or coils are positioned within the openregion 114 may also be used in embodiments of the present invention.

Reference is now made to FIG. 22 which is a schematic isometric viewillustrating an MRI probe 450 having external X, Y and Z-gradient coils,in accordance with still another preferred embodiment of the presentinvention.

The MRI probe 450 include two opposed permanent magnet assemblies 462and 464 defining an open region 414 therebetween. The permanent magnetassemblies 462 and 464 may be attached to or supported by one or moresupporting structures such as a supporting frame (not shown for the sakeof clarity of illustration) which is designed to enable access to theregion 414 and to the head of the patient 6.

An organ or body part such as the head of a patient 6 may be positionedwithin the open region 414 for imaging. The permanent magnet assemblies462 and 464 may be similar in design to the permanent magnet assemblies162 and 164 of FIG. 16 but may also be any suitably designed pair ofpermanent magnet assemblies for providing a region of substantiallyhomogenous magnetic field therebetween. The MRI probe 450 furtherincludes a transmitting RF coil 440 which includes four portions 440a,440B, 440C and 440D. the portions 440a, 440B, 440C and 440D of the RFcoil 440 are printed circuit board assemblies which are suitablyelectrically connected (connections not shown), the copper conductors(not shown) included in the printed circuit board assemblies 440a, 440B,440C and 440D are shaped in a similar way to the conductors of the RFcoil 140 of FIG. 15. However, the transmitting RF coils 140or 340 ofFIGS. 15 and 21, respectively, may also be used instead of the RF coil440. The printed circuit board assemblies 440A and 440D may also includeshim coils (not shown), however, the shim coils (not shown) may also bea pair of separate coils each disposed opposing one of the printedcircuit board assemblies 440A and 440D at a distance therefrom.

The MRI probe 450 further includes a receiving RF coil 175 and amulti-layer printed circuit assembly 472.

The multi-layer printed circuit assembly 472 is disposed underneath thepermanent magnet assemblies 462 and 464 and outside the open region 414.Thus, the region 414 may be relatively freely accessed.

The multi-layer printed circuit assembly 472 includes three printedcircuits (not shown) including a X-gradient coil, a Y-gradient coil anda Z-gradient coil. It is noted that, since the relative positioning ofthe multi-layer printed circuit assembly 472 is different than thepositioning of the multi-layer printed circuit assemblies 172 and 174 ofFIG. 16, the design of the gradient coils is adapted to suit thedifferent position of the coils relative to the direction of the mainmagnetic field. The positioning of the multi-layer printed circuitassembly 472 outside the open region 414 has the advantages of makingmore space available in the open region 414 and of reducing the loadingof the RF transmitting coil 440 by increasing the spatial separationbetween the conducting gradient coil surfaces of the multi-layer printedcircuit assembly 472 and the transmitting RF coil 440.

It is noted that while the MRI probes of the preferred embodimentsdisclosed hereinabove include a pair of opposing permanent magnetassemblies with an open region therebetween wherein an organ or bodypart is disposed in the open region between the pair of permanent magnetassemblies, other preferred embodiments of the present invention may beimplemented using a single magnet assembly.

Reference is now made to FIG. 23 which is a schematic cross section ofan MRI probe 500 having a single permanent magnet assembly, inaccordance with yet another preferred embodiment of the presentinvention.

The MRI probe 500 includes a single permanent magnet assembly 562 havinga first surface 582A and a second surface 582B opposing the firstsurface 582A. The permanent magnet assembly 562 may be constructed byusing various different designs. For example, the permanent magnetassembly 562 may be constructed from a plurality of concentric annularpermanent magnets as disclosed in detail in co-pending U.S. Pat. No.5,900,793 to Katznelson et al. However, the permanent magnet assembly562 may also be implemented using other methods and designs adapted toprovide a volume of substantially homogenous magnetic field 518extending beyond the surface 582B of the permanent magnet assembly 562.The particular design parameters of the permanent magnet assembly maydepend, inter alia, on the desired dimensions of the volume 518, thedesired intensity of the magnetic field within the volume 518 and thedistance between the volume 518 and the surface 582B.

The MRI probe 500 further includes a multi-layer printed circuitassembly 572. The multi-layer printed circuit assembly 572 is disposedopposing the surface 582A of the permanent magnet assembly 562 on theside of the permanent magnet assembly 562 which is opposite the sidefacing the volume 518. The multi-layer printed circuit assembly 572includes three printed circuits (not shown) including a X-gradient coil,a Y-gradient coil and a Z-gradient coil. The multi-layer printed circuitassembly 572 may also include a shim coil (not shown) for activeshimming of the main magnetic field.

The MRI probe 500 further includes a transmitting RF coil 540. Thetransmitting RF coil 540 is disposed between the surface 582B and thevolume 518. The MRI probe 500 further includes a receiving RF coil 575suitably connected to an RF amplifier 525 such as a low noise RFamplifier.

An organ or body part which is to be imaged, such as the head of apatient 6 can be positioned above the surface 582B and the RF coil 540such that at least part of the head 6 is positioned within the volume518.

An advantage of the MRI probe 500 is that the gradient coils and shimcoils which are included in the multi-layer printed circuit assembly 572are disposed away from the region above the surface 582B and thereforedo not restrict access of the imaged organ to the volume 518.

Another advantage of the configuration of the multi-layer printedcircuit assembly 572 within the MRI probe 500 is that the gradient andshim coils (not shown) of the multi-layer printed circuit assembly 572are positioned away from the transmitting RF coil 540 and the receivingRF coil 575 and therefore reduce the loading of the transmitting RF coil540 and of the receiving RF coil 575 by the gradient coils (not shown)within the multi-layer printed circuit assembly 572, thereby improvingimage quality.

It is noted that, other configurations of the transmitting RF coil 540are possible in which a portion of the transmitting RF coil is disposedbetween the surface 582B and the volume 518 and another portion of theRF transmitting coil is disposed between the surface 582A and themulti-layer printed circuit assembly 572.

It will be appreciated by those skilled in the art that the singlepermanent magnet assembly 562 of FIG. 23 has to be optimized to achievea predetermined volume of substantially homogenous magnetic field. Thus,three gradient coils, namely the X, Y and Z gradient coils (not shown)are needed within the multi-layer printed circuit assembly 572 toactively generate the three orthogonal magnetic field gradients requiredduring the imaging sequences.

While such optimization methods for single magnet assemblies are knownin the art, there is an alternative approach in which the single magnetassembly is designed to provide a fixed magnetic field gradient, forexample along the z-axis. For example, U.S. Pat. No. 5,390,673 toKikinis discloses a bar-like single magnet having a fixed magnetic fieldgradient along the longitudinal axis of the bar-like magnet.

The inventors of the present invention have noted that such permanentmagnet assemblies having a fixed magnetic field gradient may also beused in accordance with the present invention.

Reference is now made to FIG. 24 which is a schematic diagramillustrating an MRI probe having a fixed magnetic field gradient, inaccordance with another preferred embodiment of the present invention.

The MRI probe 600 includes a single permanent magnet assembly 662 havinga first surface 682A and a second surface 682B opposing the firstsurface 682A. The permanent magnet assembly 662 may be constructed byusing various different designs. For example, the permanent magnetassembly 662 may be constructed from as a plurality of concentricannular permanent magnets as disclosed in detail in co-pending U.S. Pat.No. 5,900,793 to Katznelson et al., now U.S. Pat. 5,900,793 wherein theexact dimensions, shapes, magnetic field strength and relativepositioning of the annular permanent magnets in the assembly aredesigned to obtain a fixed magnetic field gradient extending along theaxis 612. This fixed Z-gradient varies substantially linearly within thepredetermined volume 618 along the axis 612. The magnetic field issubstantially uniform in any plane which is included within the volume618 and is orthogonal to the axis 612 within the volume 618.

However, the permanent magnet assembly 662 may also be implemented usingother methods and designs such as the design of Kikinis. The particulardesign parameters of the permanent magnet assembly may depend, interalia, on the desired dimensions of the volume 618, the desired intensityof the magnetic field within the volume 618 and the distance between thevolume 618 and the surface 682B.

The MRI probe 600 further includes a multi-layer printed circuitassembly 672. The multi-layer printed circuit assembly 672 is disposedopposing the surface 682A of the permanent magnet assembly 662 on theside of the permanent magnet assembly 662 which is opposite the sidefacing the volume 618. In contrast to the multi-layer printed circuitassembly 572 of FIG. 23 which includes three gradient coils, themulti-layer printed circuit assembly 672 of FIG. 24 includes two printedcircuits (not shown) including an X-gradient coil, and a Y-gradientcoil. The multi-layer printed circuit assembly 672 may also include ashim coil (not shown) for active shimming of the magnetic field.

The MRI probe 600 further includes a transmitting RF coil 640, Thetransmitting RF coil 640 is disposed between the surface 682B and thevolume 618. The MRI probe 600 further includes a receiving RF coil 575suitably connected to an RF amplifier 525 such as a low noise RFamplifier.

An organ or body part which is to be imaged, such as the head of apatient 6 can be positioned above the surface 682B and the transmittingRF coil 540 such that at least part of the head 6 is positioned withinthe volume 618.

It is noted that ,while in the embodiment of FIG. 24 both the X and Ygradient coils (not shown) within the multi-layer printed circuitassembly 672 are positioned opposing the surface 682A, in otherembodiments only one of the X or Y gradient coils may be so positioned.

The advantages of the disclosed positioning of one or more of thegradient coils of the MRI probe 600 are similar to the advantagesdisclosed in detail for the MRI probe 500 hereinabove.

It is noted that, while the permanent magnet assemblies used within theMRI probes of FIGS. 4, 8-10, 14, 16-19, 22 and 23 are designed usingconcentric annular permanent magnets as disclosed in detail inco-pending U.S. Pat. No. 5,900,793 to Katznelson et al., many othertypes of magnet assemblies can be used which are within the scope of thepresent invention. For example, the annular permanent magnets used inthe construction of the permanent magnet assemblies may be concentricpolygonal annular shapes, or a plurality of elliptically shaped annulihaving two common axes passing through the foci of the individualelliptical annuli.

Additionally, other configurations of permanent magnets may be used suchas solid cube like or bar like permanent magnets or any other types ofyoked or-non yoked magnets which are constructed to avoid thedevelopment of substantial eddy currents therewithin by the gradientcoils. Such designs may use permanent magnetic materials having lowelectrical conductivity or may use magnetic and/or yoke structures whichare segmented and are attached or glued by non-electrically conductivematerials or glues. The development of eddy currents within yokestructures having high electrical conductivity may be reduced forenabling their use with the external gradient coils of the presentinvention by slotting the yoked structures with spiral or other types ofslots to reduce possible current development. Thus, the various forms ofthe external gradient positioning of the present invention may beadapted for use with differently designed magnet assemblies configuredsingly, or as opposed pairs of magnetic assemblies having an open regiontherebetween.

It is further noted that, while in the preferred embodiments of thepresent invention, the multi-layer printed circuit assemblies 52, 54,72, 74, 64, 78, 172, 174, 272, 274, 200, 202, 472 and 572 includetherewithin hollow conduits (not Shown) for flowing a coolant fluid suchas water therein, in other preferred embodiments of the presentembodiments, the multi-layer printed circuit assemblies may be devoid ofsuch channels.

It is still further noted that, in accordance with yet other preferredembodiments of the present invention, screening devices such asconducting metal mesh or grid may be inserted between various componentsof the MRI probes for improving RF screening. For example, in the MRIprobe 150 of FIGS. 16 and 17, a suitable circular copper mesh piece (notshown) of a diameter similar to the diameter of the multi-layer printedcircuit assembly 172 may be disposed between the surface 182A and themulti-layer printed circuit assembly 172, while another suitablecircular copper mesh piece (not shown) of a diameter similar to thediameter of the multi-layer printed circuit assembly 174 may be disposedbetween the surface 184B and the multi-layer printed circuit assembly174. Similarly, pieces of suitable copper mesh (not shown) may be usedfor screening the entire surface of the permanent magnet assembly 162except the surface 182B thereof, and the entire surface of the permanentmagnet assembly 164 except the surface 184A thereof.

It is also noted that for the sake of clarity of illustration in many ofthe drawing Figures, the transmitting and receiving RF amplifiers, andthe precise electrical connections between the portions of the RF coilsare not shown.

It is further noted that, preferably, in all the embodiments of the MRIprobes illustrated in FIGS. 4, 8-10, and 16-19, all the correspondingpairs of the gradient coils and shim coils of the MRI probe areelectrically connected in series (the connections are not shown for thesake of clarity of illustration). For example, the Z-gradient amplifier(not shown) of the MRI probe 150 of FIG. 16 is electrically connected tothe current input terminal (not shown) of the z-gradient coil (notshown) included within the multi-layer printed circuit assembly 172, thecurrent output terminal (not shown) of the z-gradient coil of themulti-layer printed circuit assembly 172 is electrically connected tothe current input terminal (not shown) of the z-gradient coil (notshown) included within the multi-layer printed circuit assembly 174, andthe current output terminal (not shown) of the z-gradient coil of themulti-layer printed circuit assembly 174 is electrically connected tothe Z-gradient amplifier, completing the circuit. Thus the z-gradientamplifier energizes both of the complementary z-gradient coils of theopposing multi-layer printed circuit assemblies 172 and 174,simultaneously. A similar in-series electrical connection scheme is usedfor the pairs of complementary x-gradient coils (not shown) y-gradients(not shown) and Shim coils (not shown). However, other methods ofconnection of the complementary pairs of gradient and shim coils mayalso be used, such as the use of pairs of amplifiers (not shown), eachof which activates one coil of the complementary pairs of coils.

While embodiments of the present invention have been described so as toenable one skilled in the art to practice the present invention, thepreceding description is intended to be exemplary and should not beconstrued as limiting the scope of the invention.

What is claimed is:
 1. Electromagnetic apparatus for use in an MRIdevice, the apparatus comprising:a permanent magnet assembly having afirst surface defining a first side of said permanent magnet assemblyand a second surface defining a second side of said permanent magnetassembly opposed to said first side, for producing a predeterminedvolume of substantially uniform magnetic field extending in a firstdirection beyond said first surface; an energizable transmitting RF coilfor producing an RF electromagnetic field within said volume, at least aportion of said RF coil is positioned adjacent said first surface ofsaid permanent magnet assembly; an energizable z-gradient coil forproducing a magnetic field gradient extending within said volume in saidfirst direction parallel to a first axis; an energizable x-gradient coilfor producing a magnetic field gradient extending within said volumeparallel to a second axis orthogonal to said first axis; and anenergizable y-gradient coil for producing a magnetic field gradientextending within said volume parallel to a third axis orthogonal to saidfirst axis and to said second axis, wherein at least one of saidx-gradient coil, y-gradient coil and z-gradient coil is positionedopposing said second surface of said permanent magnet assembly.
 2. Theapparatus according to claim 1 wherein at least a portion of saidtransmitting RF coil is positioned on said second side of said permanentmagnet assembly opposing said second surface of said permanent magnetassembly to improve the efficiency of said transmitting RF coil.
 3. Theapparatus according to claim 1 further including at least oneenergizable shim coil for improving the homogeneity of saidsubstantially uniform magnetic field.
 4. The apparatus according toclaim 3 wherein said at least one shim coil is a substantially planarcoil opposing said second surface of said permanent magnet assembly. 5.The apparatus according to claim 1 wherein said x-gradient coil, saidy-gradient coil and said z-gradient coil are substantially planarprinted circuits assembled within a substantially planar multi-layerprinted circuit assembly, said multi-layer printed circuit assembly isdisposed on said second side of said permanent magnet assembly facingsaid second surface.
 6. The apparatus according to claim 5 wherein saidmulti-layer printed circuit assembly further includes at least oneenergizable shim coil, said at least one shim coil is a substantiallyplanar printed circuit.
 7. The apparatus according to claim 1 furtherincluding at least one receiving RF coil positioned on said first sideof said permanent magnet assembly and placeable adjacent to an organ orbody part to be imaged using said apparatus.
 8. The apparatus accordingto claim 7 wherein said transmitting RF coil is a linearly polarizing RFcoil.
 9. The apparatus according to claim 7 wherein said transmitting RFcoil is a circularly polarizing RF coil.
 10. The apparatus according toclaim 1 wherein said permanent magnet assembly comprises:a first annularpermanent magnet having an upper and a lower surface thereof, said uppersurface of said first annular permanent magnet being of a first magneticpolarity and said lower surface of said first annular permanent magnetbeing of a second magnetic polarity, said first annular permanent magnethaving an inside diameter, said first permanent magnet having at least aportion of said upper surface of said first annular magnet lying in afirst plane and providing a first magnetic field in said predeterminedvolume, said first magnetic field having a zero rate of change in saidfirst direction at a first point; at least a second annular permanentmagnet having an upper and a lower surface thereof, said upper surfaceof said at least second annular permanent magnet being of said firstmagnetic polarity and said lower surface of said at least second annularpermanent magnet being of said second magnetic polarity, said at leastsecond annular permanent magnet having an outside diameter which issmaller than said inside diameter of said first annular permanentmagnet, said at least second annular permanent magnet providing a secondmagnetic field; and low permeability material interconnecting said firstannular permanent magnet with said at least second annular permanentmagnet, so that at least a portion of said upper surface of said atleast second annular permanent magnet is in a second plane spaced fromsaid first plane, whereby said second magnetic field is superimposedupon said first magnetic field, in said predetermined volume, having azero rate of change in said first direction at a second point differentfrom said first point.
 11. The apparatus according to claim 10 whereinsaid first axis passes through the center points of said first annularpermanent magnet and said at least second annular permanent magnet. 12.The apparatus according to claim 10 wherein said z-gradient coil is anextended gradient coil concentrically disposed between said firstannular permanent magnet and said at least second annular permanentmagnet, said z-gradient coil has a longitudinal axis coincident withsaid first axis.
 13. The apparatus according to claim 10 wherein saidfirst annular permanent magnet and said at least second annularpermanent magnet are rare-earth permanent magnets.
 14. The apparatusaccording to claim 13 wherein said rare-earth permanent magnets areneodimium-iron-boron alloy permanent magnets.
 15. The apparatusaccording to claim 10 wherein at least one of said first annularpermanent magnet and said at least second annular permanent magnetcomprises a plurality of segments attached to adjacent segments using anelectrically non-conductive adhesive.
 16. The apparatus according toclaim 15 wherein said segments are equiangular segments.
 17. Theapparatus according to claim 15 wherein said segments have a trapezoidalcross-section in a plane orthogonal to said first direction. 18.Electromagnetic apparatus for use in an MRI device, the apparatuscomprising:a permanent magnet assembly having a first surface and asecond surface for producing a predetermined volume having a magneticfield varying substantially linearly along a first axis, said volumeextending in a first direction beyond said first surface along saidfirst axis, said magnetic field being substantially uniform in any planeincluded within said predetermined volume and orthogonal to said firstdirection within said predetermined volume; an energizable transmittingRF coil for transmitting RF radiation, said RF coil having at least oneportion thereof positioned opposing said first surface of said permanentmagnet assembly; an energizable x-gradient coil for producing a magneticfield gradient along a second axis orthogonal to said first axis; and anenergizable y-gradient coil for producing a magnetic field gradientalong a third axis orthogonal to said first axis and to said secondaxis, wherein at least one of said x-gradient coil and y-gradient coilis positioned opposing said second surface of said permanent magnetassembly.
 19. The apparatus according to claim 18 further including atleast one receiving RF coil positioned on said first side of saidpermanent magnet assembly and placeable adjacent to an organ or bodypart to be imaged using said apparatus.
 20. A method for constructing anelectromagnetic apparatus for use in an MRI device, the methodcomprising the steps of:providing a permanent magnet assembly having atleast a first surface defining a first side of said permanent magnetassembly and a second surface defining a second side of said permanentmagnet assembly opposed to said first side, for producing apredetermined volume of substantially uniform magnetic field extendingin a first direction beyond said first surface; providing an energizabletransmitting RF coil for producing an RF electromagnetic field withinsaid volume; positioning at least a portion of said transmitting RF coiladjacent said first surface of said permanent magnet assembly; providingat least one receiving RF coil placeable adjacent to an organ or bodypart to be imaged for receiving RF signals from said organ or body part;providing an energizable z-gradient coil for producing a magnetic fieldgradient extending within said volume in said first direction parallelto a first axis; providing an energizable x-gradient coil for producinga magnetic field gradient extending within said volume parallel to asecond axis orthogonal to said first axis; providing an energizabley-gradient coil for producing a magnetic field gradient extending withinsaid volume parallel to a third axis orthogonal to said first axis andto said second axis; and positioning at least one of said x-gradientcoil, y-gradient coil and z-gradient coil opposite said second surfaceof said permanent magnet assembly for reducing the loading of saidtransmitting RF coil and said at least one receiving RF coil by said atleast one of said x-gradient coil, y-gradient coil and z-gradient coil.21. A method for constructing electromagnetic apparatus for use in anMRI device, the method comprising the steps of:providing a permanentmagnet assembly having a first surface and a second surface forproducing a predetermined volume having a magnetic field varyingsubstantially linearly along a first axis, said volume extending in afirst direction beyond said first surface along said first axis, saidmagnetic field being substantially uniform in any plane included withinsaid predetermined volume and orthogonal to said first direction withinsaid predetermined volume; providing an energizable transmitting RF coilfor transmitting RF radiation; positioning said transmitting RF coilsuch that at least one portion thereof opposes said first surface ofsaid permanent magnet assembly; providing at least one receiving RF coilplaceable adjacent to an organ or body part to be imaged for receivingRF signals from said organ or body part; providing an energizablex-gradient coil for producing a magnetic field gradient along a secondaxis orthogonal to said first axis; providing an energizable y-gradientcoil for producing a magnetic field gradient along a third axisorthogonal to said first axis and to said second axis; and positioningat least one of said x-gradient coil and y-gradient coil opposite saidsecond surface of said permanent magnet assembly for reducing theloading of said transmitting RF coil and said at least one receiving RFcoil by said at least one of said x-gradient coil and y-gradient coil.